1. Technical Field
In general, the present disclosure relates to nuclear medical imaging. More particularly, the disclosure relates to Positron Emission Tomography (PET) and Single Photon Emission Computed Tomography (SPECT) compensating for changes to the gain of a photo detector due, inter alia, temperature variation.
2. General Background of the Invention
Nuclear medicine is a unique specialty wherein radiation emission is used to acquire images that show the function and physiology of organs, bones or tissues of the body. The technique of acquiring nuclear medicine images entails first introducing radiopharmaceuticals into the body—by either injection or ingestion. These radiopharmaceuticals are attracted to specific organs, bones, or tissues of interest. The radiopharmaceuticals produce gamma photon emissions, which emanate from the body and are then captured by a scintillation crystal. The interaction of the gamma photons with the scintillation crystal produces flashes of light or electromagnetic radiation in a different spectrum, which are referred to as “scintillation events.” Scintillation events can be detected by an array of photo detectors (such as photomultiplier tubes (PMT) or avalanche photodiodes (APD)), and the event pulses are processed and stored. The stored data can be processed in the reconstruction system and a volume image illustrating the distribution of the radioisotopes in the body is reconstructed. The tracer distribution image can be overlayed with morphological images produced by other imaging modalities such as CT or MR. For example, CT and MR system can be integrated with other nuclear imaging systems to create a hybrid to yield more information to a health practitioner.
One particular nuclear medicine imaging technique is known as positron emission tomography, or PET. PET is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. The measurement of tissue concentration using a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from a positron annihilation or coincidence event. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by a coincidence event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Coincidence events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors; i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line-of-response (LOR) along which the coincidence event has occurred. An example of a PET method and apparatus is described in U.S. Pat. No. 6,858,847, which patent is incorporated herein by reference in its entirety.
The gamma photons detected by the scintillator can be converted into scintillation photons, which can be detected by the photo sensors (e.g. PMT or APD). The photo sensor can convert the light into a current, which generates electric pulses. The pulses are proportional to the energy of the impinging gamma photons detected by the scintillator and can be processed and stored by the system front-end electronics. The processed pulses, also called events, can be processed into an energy histogram, i.e. energy spectrum (FIG. 2). The dynamic range of the front-end electronic processing will determine the number of bins (typically 128 or 256) in the energy spectrum. Based on the photo-peak position 23, representing the gamma energy of the radioisotope of the tracer (e.g., 511 keV for PET annihilation gamma radiation), the spectrum of energy received by the photo sensors can be calibrated and the bin width calculated (e.g., bin width=256/511 keV≈2 keV). In order to discriminate scattered radiation from unscattered “true” events, a lower-level discriminator (LLD) 21 and upper-level discriminator (ULD) 22 are defined around the energy peak 23. Only events falling within the window of the LLD and ULD will be qualified and processed for reconstruction. The stability of the photo peak position within the energy spectrum is of utmost importance and will directly relate to image quality of the reconstructed tracer distribution.